The present invention relates to gamma cameras and more specifically to a method and apparatus for use with a scintillation camera for correcting distortions in images acquired simultaneously at two different energy levels.
In emission imaging, examinations are carried out by injecting a dilution marker comprising a compound labeled with a radiopharmaceutical into the body of a patient to be examined. A radiopharmaceutical is a substance that emits photons at one or more energy levels. By choosing a compound that will accumulate in an organ to be imaged (hereinafter an organ of interest), compound concentration, and hence radiopharmaceutical concentration, can be substantially limited to an organ. Typically a radiopharmaceutical that emits photons at approximately a single known energy level is chosen. An energy range which approximates the known energy level will be referred to as the marker range.
While moving through a patient's blood stream the marker, including the radiopharmaceutical, becomes concentrated in the organ of interest. By measuring the number of photons emitted from the organ of interest which are within the marker range, organ characteristics, including irregularities, can be identified.
To measure the number of emitted photons, one or more planar gamma cameras are used. After a marker has become concentrated within an organ of interest, a camera is positioned at an imaging angle with respect to the organ of interest such that the organ is positioned within the camera's field of view (FOV). The camera is designed to detect photons traveling along preferred paths within the FOV.
A gamma camera consists of a collimator, a scintillation crystal, a plurality of photomultiplier tubes (PMTs) and a camera processor. The collimator typically includes a rectangular lead block having a width dimension and a length dimension which together define the FOV. The collimator block forms tiny holes which pass therethrough defining preferred photon paths which are unidirectional and perpendicular to the length of the collimator. The collimator blocks emissions toward the crystal along non-preferred paths.
The scintillation crystal is positioned adjacent the collimator on a side opposite the FOV and has an impact surface and an oppositely facing emitter surface. The impact surface defines a two dimensional imaging area A having a length L and a width W. Photons which pass through the collimator impact and are absorbed by the impact surface at impact points, each point having actual impact coordinates X and Y along length L and width W dimensions. The crystal emitter surface emits light from an emitter point adjacent the impact point each time a photon is absorbed. The amount of light emitted depends on the absorbed photon's energy level.
The PMTs are arranged in a two dimensional array which is positioned adjacent the emitter surface. Light emitted by the crystal is detected by the PMTs. Each PMT which detects light generates an analog intensity signal which is proportional to the amount of light detected. When a single photon is absorbed by the crystal, the emitted light is typically absorbed by several different PMTs such that several PMTs generate intensity signals simultaneously. For the purposes of this explanation all intensity signals caused by a single photon will be collectively referred to as a signal set.
The processor receives each signal set and performs a plurality of calculations on each signal set to determine two characteristics of the corresponding photon. First, the processor combines the intensity signals of each signal set to identify the energy level Z of a corresponding photon. Photons having energies within the marker range will be referred to as events. Only signals corresponding to events are used for imaging. Second, the processor performs a series of calculations in an effort to determine precise impact coordinates X and Y of each event. Hereinafter, actual impact coordinates will be referred to as X and Y while calculated coordinates will be referred to as Xc and Yc. Once coordinates X and Y of all events have been identified, the processor uses the coordinates of many events (typically millions) to create an image of the organ of interest.
Unfortunately, images formed from calculated impact coordinates Xc and Yc are often spatially distorted. Spatial distortion occurs when calculated coordinates Xc and Yc diverge from actual event coordinates X and Y. This distortion occurs because calculated coordinates Xc and Yc are dependent upon, and are a function of, the spatial relationship between a photon's impact point and adjacent PMTs. The effect of this type of distortion is that the number of events perceived in the central areas of many PMTs is greater than the actual number of events which occur in those areas while the number of events perceived in edge areas of many PMTs is less than the actual number of events which occur. This incorrect event concentration causes images generated using the event information to have a non-uniform appearance and therefore limits their diagnostic usefulness.
To compensate for PMT spatial distortion, the industry has developed spatial coordinate correction methods usually implemented in software. According to these methods, after event coordinates Xc and Yc have been calculated, the coordinates are then shifted as a function of coordinate correction factors to compensate for incorrect event concentrations along both X and Y coordinate axis. The shifts generate modified coordinates Xm and Ym for each photon wherein the modified coordinates are nearly identical to actual coordinates X and Y.
To identify correction factors, a radiation image having a known pattern and known energy such as a grid of parallel lines or points (typically formed using a radio-opaque grid), is directed at a camera. Xc and Yc coordinates are then computed and compared with known coordinates X and Y which correspond to the grid pattern. The comparison yields correction factors which, when applied to calculated coordinates Xc and Yc, generate modified coordinates Xm and Ym which are identical to actual coordinates X and Y.
In medical imaging, it is sometimes desirable to generate two nuclear images simultaneously using first and second radiopharmaceuticals, the first radiopharmaceutical emitting photons at a first energy and the second radiopharmaceutical emitting photons at a second energy different than the first. For example, photons at the first energy might be better suited for identifying scarred tissue while photons at the second energy might be better suited for imaging blood flow or some other biological process or tissue. The quantity of light produced by the crystal upon photon absorption is approximately linearly proportional to an absorbed photons energy. For example, a photon of 511 keV would produce approximately 3.6 times the quantity of light produced by a photon of 140 keV. The processor can identify each photon energy and separate absorbed photons into two different groups, each group including photons at one of the two energy levels. Then, the processor can construct two different images using modified impact coordinates Xm and Ym corresponding to the two different photon groups.
In addition to being dependent on PMT location and construction, spatial coordinate correction factors are also dependent on photon energy. This is because photon energy affects operation of the processor's electronic circuitry in a non-linear fashion. For example, where a photon of 511 keV is absorbed at coordinates X and Y, the processor might generate calculated coordinates Xca ands Yca whereas, where a photon of 140 keV is absorbed at coordinates X and Y, the processor might generate calculated coordinates Xcb and Ycb which are slightly different than coordinates Xca and Yca.
The most obvious solution to correct for non-linearities due to different photon energies is to generate one set of spatial coordinate correction factors for a first energy and a second set of spatial coordinate correction factors for a second energy and then apply the first and second factor sets to calculated impact coordinates corresponding to photons at the first and second energies, respectively. One such method has been described in U.S. Pat. No. 5,345,082 entitled "Scintillation Camera Utilizing Energy Dependent Linearity Correction" which issued on Sep. 6, 1994 which is incorporated herein by reference. According to that patent, during imaging and after a processor detects a photon event, the processor generates energy signal Z corresponding to the event, compares signal Z with energy signal ranges corresponding to different radiopharmaceutical energy levels to identify an energy range including signal Z, computes calculated coordinates Xc and Yc, retrieves a correction factor set corresponding to the energy range, identifies correction factors within the retrieved set which correspond to calculated coordinates Xc and Yc and modifies coordinates Xc and Yc according to the correction factors. This entire process is performed during data acquisition.
While this solution to the energy distortion problem can facilitate distortion correction, this solution cannot be implemented by every gamma camera processor. This is because many camera processors were designed to correct for spatial distortion corresponding to photons at a single energy level and therefore cannot support distortion correction corresponding to events at more than one energy without hardware modifications. In these cases, while processor hardware might be able to separate events into first and second energy ranges, the processor can only modify calculated coordinates according to a single correction factor set corresponding to the first energy range. Thus, while modified coordinates corresponding to events within the first energy range will be correct, modified coordinates corresponding to events within the second energy range will be incorrect and an image formed using the second set of modified coordinates will be will be of lower quality for diagnostic purposes.
In addition, this solution requires an extremely complex and hence expensive processor to perform all processor functions during data acquisition. One way to implement this solution less expensively might be to store sets of acquired data during acquisition for later "off-line" processing. If the position coordinates are reduced to, say 7 or 8 bits as is usual in nuclear medicine imaging, some position information is lost which renders subsequent correction less accurate. This problem is worst for completely uncorrected images where the distortion may be relatively large.
Therefore, it would be advantageous to have a universal method for correcting calculated event coordinates for events at two or more energy levels which could be implemented using essentially any camera processor without modification to existing hardware and without reduction in data accuracy.